Wireless Midfield Systems and Methods

ABSTRACT

Implantable devices and/or sensors can be wirelessly powered by controlling and propagating electromagnetic waves in a patient&#39;s tissue. Such implantable devices/sensors can be implanted at target locations in a patient, to stimulate areas such as the heart, brain, spinal cord, or muscle tissue, and/or to sense biological, physiological, chemical attributes of the blood, tissue, and other patient parameters. The propagating electromagnetic waves can be generated with sub-wavelength structures configured to manipulate evanescent fields outside of tissue to generate the propagating waves inside the tissue. Methods of use are also described.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a Continuation of U.S. patent application Ser. No.15/022,374, filed Mar. 16, 2016, which application is a U.S. NationalStage Filing under 35 U.S.C. 371 from International Application No.PCT/US2014/055885, filed Sep. 16, 2014, which application claims thebenefit of U.S. Provisional Appln. No. 61/878,436, filed Sep. 16, 2013,titled “Multi-Element Coupler”, and U.S. Provisional Appln. No.61/913,164, filed Dec. 6, 2013, titled “Power Management and Conversionfor Medical Implants”, all of which are incorporated herein byreference.

INCORPORATION BY REFERENCE

All publications and patent applications mentioned in this specificationare herein incorporated by reference to the same extent as if eachindividual publication or patent application was specifically andindividually indicated to be incorporated by reference.

FIELD

This disclosure is related generally to wireless power transfer. Morespecifically, this disclosure relates to delivering wireless powerthrough tissue into a device implanted in a human or animal.

BACKGROUND

Systems and methods that supply power without electrical wiring aresometimes referred to as wireless energy transmission (WET). Wirelessenergy transmission greatly expands the types of applications forelectrically powered devices. Implantable medical devices typicallyrequire an internal power source able to supply adequate power for thereasonable lifetime of the device or an electrical cable that traversesthe skin.

More recently there has been an emphasis on systems that supply power toan implanted device without using transcutaneous wiring, sometimesreferred to as a Transcutaneous Energy Transfer System (TETS).Frequently energy transfer is accomplished using two magneticallycoupled coils set up like a transformer so power is transferredmagnetically across the skin. Conventional systems are relativelysensitive to variations in position and alignment of the coils,typically requiring the coils to be physically close together and wellaligned.

Existing systems that transmit power wirelessly based on magnetic fieldstypically operate in the near-field only, where the separation of thetransmitter and receiver coils is less than or equal to the dimension ofthe coils.

Wireless powering has long been of interest for enhancing the functionof implantable electronics, beginning in the early 1960's withexperiments in transporting electromagnetic energy across the chestwall. Drawing conceptually on schemes for transferring power over airthrough objects coupled in the near-field, early manifestations involvedbulky coils tether to vacuum tube power supplies or battery cells thatposed severe challenges for long-term operation in the body. Advances insemiconductor technology have since enabled sophisticated devices thatincorporate sensing and stimulation capabilities within cellular-scaledimensions. Nearly all existing systems, however, continue to requirelarge structures for energy storage or harvesting, often severalcentimeters in the largest dimension with overall size, weight, andefficiency characteristics that constrain opportunities for integrationinto the body.

Near-field approaches rely on strong coupling occurring between objectswith matched electrical characteristics, such as resonances andimpedances. These near-field approaches do not generalize easily togeometries with extreme size asymmetry, while far-field transfer islimited by absorption over surfaces of the body.

The present disclosure describes methods and apparatus for wirelesspower transfer that overcome the limitations of previous wireless powertransfer methods. The present disclosure provides a mid-field approachin which both evanescent and radiative components of a structure arecoupled to modes in tissue that transport energy continuously away fromthe source. Interference resulting from phase differences between thesecomponents affords additional opportunity for spatially focused anddynamically adjustable field patterns inside tissue. The level ofperformance obtainable from the approach described in this disclosurecan exceed requirements for advanced monitoring and control capabilitiesfor applications in medicine, neuroscience, or human-machine interfaces.

SUMMARY

In one embodiment, a wireless power system is provided, comprising anexternal module having one or more sub-wavelength structures configuredto transmit wireless power by manipulating evanescent fields outside oftissue to generate propagating fields inside the patient's tissue andthereby generate a spatially focused field in tissue, and an implantablemodule configured to receive the wireless power from the externalmodule, the implantable module including at least one sensor orstimulator configured to sense a parameter of the tissue or applystimulation to the tissue.

In some embodiments, the at least one sensor is selected from the groupconsisting of a thermal sensor, a chemical sensor, a pressure sensor,and oxygen sensor, a PH sensor, a flow sensor, an electrical sensor, astrain sensor, a magnetic sensor, and an imaging sensor.

In other embodiments, the at least one stimulator is selected from thegroup consisting of an electrical stimulator, an optical stimulator, achemical stimulator, and a mechanical stimulator.

In one embodiment, the implantable device comprises a modular designthat allows interchangeable sensors and/or stimulators.

In some embodiments, the one or more sub-wavelength structures areselected from the group consisting of a patch, a PIFA, a slot in aground plane, a cross slot in a ground plane, an aperture coupledcircular slot in a ground plane, and a half slot in a ground plane.

In another embodiment, the external module further comprises one or moreexcitation ports coupled to the one or more sub-wavelength structures,at least one voltage source coupled to the one or more excitation ports,and a controller configured to adjust a phase and/or an amplitudedelivered to the one or more sub-wavelength structures to adjust aposition of a focal point of the spatially focused field in the tissue.

In one embodiment, the controller is configured to detect a power levelof received wireless energy from the implanted module, and is configuredto provide feedback to automatically adjust the position of the focalpoint to optimize wireless power transmission.

In another embodiment, the implantable module is configured to beimplanted on, in, or near a heart to apply leadless pacing to the heart.

In some embodiments, the implantable module is configured to beimplanted on, in, or near a brain to apply deep brain stimulation to thebrain. In another embodiment, the implantable module is configured to beimplanted on, in, or near a spinal cord to apply stimulation to thespinal cord. In yet another embodiment, the implantable module isconfigured to be implanted on, in, or near a muscular tissue of thetongue to apply stimulation to the tongue to treat obstructive sleepapnea.

A method of providing therapy to a patient is provided, comprisingimplanting a wireless power receiving in the patient, transmitting amid-field propagating wave to the wireless power receiving module topower the module, sensing a parameter of the patient with the wirelesspower receiving module, and providing a therapy to the patient with thewireless power receiving module based on the sensed parameter.

In some embodiments, the transmitting step further comprisesmanipulating evanescent fields outside of the patient's tissue togenerate propagating fields inside the patient's tissue and therebygenerate a spatially focused field in the tissue. A method of cardiacpacing in a patient is also provided, comprising implanting a wirelesspower receiving module in, on, or near a heart, transmitting a mid-fieldpropagating wave to the wireless power receiving module to power themodule, sensing a parameter of the heart with the wireless powerreceiving module, and providing electrical pacing to the heart with thewireless power receiving module based on the sensed parameter.

In some embodiments, the transmitting step further comprisesmanipulating evanescent fields outside of the patient's tissue togenerate propagating fields inside the patient's tissue and therebygenerate a spatially focused field in the tissue.

A method of deep brain stimulation is also provided, comprisingimplanting a wireless power receiving module in, on, or near a brain,transmitting a mid-field propagating wave to the wireless powerreceiving module to power the module, sensing a parameter of the brainwith the wireless power receiving module, and providing stimulation tothe brain with the wireless power receiving module based on the sensedparameter.

In some embodiments, the transmitting step further comprisesmanipulating evanescent fields outside of the patient's tissue togenerate propagating fields inside the patient's tissue and therebygenerate a spatially focused field in the tissue.

A method of stimulating tissue is provided, comprising implanting awireless power receiving module into tissue, transmitting a mid-fieldpropagating wave to the wireless power receiving module to power themodule, sensing a parameter of the tissue with the wireless powerreceiving module, and providing stimulation to the tissue with thewireless power receiving module based on the sensed parameter.

In some embodiments, the transmitting step further comprisesmanipulating evanescent fields outside of the patient's tissue togenerate propagating fields inside the patient's tissue and therebygenerate a spatially focused field in the tissue.

In another embodiment, the method further comprises adjusting a focalpoint of the propagating wave to optimize wireless power transmission tothe module.

In another embodiment, the transmitting step comprises transmitting thewave with a sub-wavelength structure that produces a magnetic fieldperpendicular to the wave and parallel to a tissue interface.

An apparatus configured to transfer wireless power through tissue isprovided, comprising a substrate, at least one sub-wavelength structuredisposed on the substrate, at least one radio-frequency port coupled tothe at least one sub-wavelength structure, a voltage or current sourcecoupled to the at least one radio-frequency port, and a controllerconfigured to manage excitation of the at least one radio-frequency portand sub-wavelength structure with the voltage or current source tomanipulate evanescent fields outside of tissue to generate propagatingfields inside the tissue and thereby generate a spatially focused fieldin the tissue.

In some embodiments, each of the at least one sub-wavelength structureis coupled to a respective independent radio-frequency port.

An apparatus configured to transfer wireless power through tissue isalso provided, comprising a plurality of sub-wavelength structuresconfigured and arranged to generate propagating fields inside tissue andthereby generate a spatially adaptable electromagnetic field in thetissue, a plurality of independent feed ports configured and arranged toindividually excite a respective one of the plurality of sub-wavelengthstructures thereby generating the spatially adaptable electromagneticfield, and a controller configured to redistribute a peak surfaceelectromagnetic field to increase an allowable radio frequency outputpower.

In some embodiments, the plurality of sub-wavelength structures arefurther configured and arranged to generate an adaptive steering fieldin tissue.

In other embodiments, the spatially focusing and adaptive steeringfield/signal has a frequency between 300 MHz and 3000 MHz.

BRIEF DESCRIPTION OF THE DRAWINGS

The novel features of the invention are set forth with particularity inthe claims that follow. A better understanding of the features andadvantages of the present invention will be obtained by reference to thefollowing detailed description that sets forth illustrative embodiments,in which the principles of the invention are utilized, and theaccompanying drawings of which:

FIGS. 1A-1N show various embodiments of an external wireless powertransmitting module.

FIG. 2 shows the magnetic field that results from a conventionalinductively coupled loop source.

FIG. 3A shows the magnetic field that results from a patchsub-wavelength structure.

FIG. 3B shows the magnetic field that results from a PIFA sub-wavelengthstructure.

FIG. 3C shows the magnetic field that results from an aperture coupledcircular slot sub-wavelength structure.

FIG. 3D shows the magnetic field that results from a cross slotsub-wavelength structure.

FIG. 3E shows the magnetic field that results from a half slotsub-wavelength structure.

FIG. 4A shows an implanted device in a human patient being wirelesslypowered by a mid-field propagating wave technique.

FIGS. 4B-4C show wireless power transmission with an inductively coupledapproach (FIG. 4B) and a mid-field propagating wave approach (FIG. 4C).

FIGS. 5A-5B are schematic diagrams of architectures for a controller ofthe external module of FIGS. 1A-1N.

FIG. 6 shows one embodiment of an implanted device configured to receivewireless power from the external module of FIGS. 1A-1N.

FIGS. 7A-7C show embodiments of architectures for a controller of theimplanted device of FIG. 6.

FIG. 8 shows several arrangements of sub-wavelength structures andrepresentative tissue.

FIGS. 9A-9E show a multilayer model of power transfer.

FIGS. 10A-10G show midfield power transfer realization with a patternedmetal plate.

FIG. 11 shows a reflected power spectrum as a function of frequency.

FIG. 12 shows a power transfer efficiency chart.

FIGS. 13A-13D show an example of device movement.

FIGS. 14A and 14B show schematics of a power measurement probe.

FIGS. 15A and 15B show slot-array control components.

FIG. 16 shows a table with fabrication process and power consumptionvalues.

DETAILED DESCRIPTION

Implantable devices and/or sensors can be wirelessly powered bycontrolling and propagating electromagnetic waves in a patient's tissue.The implantable devices can be implanted in humans or in other animalssuch as pets, livestock, or laboratory animals such as mice, rates, andother rodents. Such implantable devices/sensors can be implanted attarget locations in a patient, as non-limiting examples, to stimulateareas such as the heart, and/or to sense biological, physiological,chemical attributes of the blood, tissue, and other patient aspects.Difficulties in achieving wireless power transfer can occur in themismatch between the size of the implantable devices/sensors and thepower transfer source, the depth of the devices/sensors in a patient,and additionally the spatial arrangement of the devices/sensors relativeto the power transfer source.

Various aspects of the present disclosure are directed towardapparatuses or methods as exemplified or supported by aspects of theabove noted description/embodiments, as well as thedescription/embodiments of the attached appendices. For instance,certain embodiments of the present disclosure are directed tomanipulation of evanescent fields outside a patient's tissue toexcite/control propagating fields inside the patient's tissue andthereby generate a spatially focusing and adaptive steering field/signalin the tissue.

Each of the sub-wavelength structures described above can be connectedto a respective port in order to manipulate evanescent fields toexcite/control propagating fields inside a patient's tissue. Thesepropagating fields can be further manipulated to generate a spatiallyfocusing and adaptive steering field/signal in the tissue. Anysub-wavelength structure that yields transverse magnetic fieldsdominating near the source, will minimize the tissue heating effect.These sub-wavelength structures can be configured to generate a magneticnear field that is in parallel with the tissue interface, and that isperpendicular with the propagating wave that transmits wireless energy.In certain embodiments, as shown above, an arrangement can include one,two, three, or four or more sub-wavelength structures are used tomanipulate the evanescent fields. In other embodiments, two or more ofthe arrangements shown above can be combined such that even moresub-wavelength structures (e.g., six, eight, twelve) are used tomanipulate the evanescent fields.

In certain embodiments, an arrangement can include two, three, four, ormore sub-wavelength structures that can be configured to manipulate theevanescent fields. In other embodiments, two or more of the arrangementsshown above can be combined such that even more sub-wavelengthstructures (e.g., six, eight, twelve, or more) are used to manipulatethe evanescent fields.

Various aspects of the present disclosure include apparatus and methodsdirected to multiple sub-wavelength structures configured to generate aspatially adaptable electromagnetic field/signal (e.g., a midfieldelectromagnetic field) in a patient's tissue. The sub-wavelengthstructures can each be connected to an independent feed port thatindividually excites a respective one of the sub-wavelength structures,thereby generating the spatially adaptable electromagnetic field/signal.The independent feed ports and the sub-wavelength structures (such aspatterned coils, conductors, and the like) are adapted to redistributethe peak surface electromagnetic fields in order to increase theallowable radio frequency output power in accordance with regulationsfrom the apparatus.

In certain embodiments, the sub-wavelength structures manipulateevanescent fields to excite/control propagating fields and therebygenerate a spatially focusing and adaptive steering field/signal intissue.

Various aspects of the present disclosure include apparatus and methodsdirected to multiple sub-wavelength structures that generate and receivea spatially adaptable electromagnetic field/signal, which may include apower signal and a communication data signal. Additionally, aspects ofthe present disclosure may also include multiple sub-wavelengthstructures that generate a spatially adaptable electromagneticfield/signal and to provide and receive a spatially adaptableelectromagnetic signal in multiple frequency bands.

Certain aspects of the present disclosure are also directed towardimplantable devices that receive power transmitted via thesub-wavelength structures that transmit a spatially adaptableelectromagnetic field. The implantable device, consistent with variousaspects of the present disclosure, can be a size such that the device isdeliverable via a catheter, cannula, or a needle. Additionally, theimplantable device(s) can include a coil that receives the energy fromthe spatially adaptable electromagnetic field. In such an embodiment,the spatially adaptable electromagnetic field/signal is received asmagnetization due to current in the coil. Further, the implantabledevices can also include, in certain instances, a multi-turn coil thatreceives the spatially adaptable electromagnetic signal, rectifyingcircuitry that converts the spatially adaptable electromagnetic signalusing AC-DC power conversion, and control circuitry to regulate pulseamplitudes, duration, and frequency.

Additionally, in certain embodiments, the sub-wavelength structures,consistent with various aspects of the present disclosure, adjust anoperating frequency of the spatially adaptable electromagnetic signal toadjust the power of the implantable device or sensor. In someembodiments, the spatially adaptable electromagnetic signal can havefrequency between 300 MHz and 3000 MHz.

Various aspects of the present disclosure are directed toward poweringof one or more active implantable sensors or devices using a singlepower source. The types of implantable devices/sensors that can bepowered using the single power source, consistent with various aspectsof the present disclosure, are numerous. For instance, the implantabledevices can be used for muscular stimulation, stimulation/sensing toregulate a patient's heart beat, multisite deep brain stimulation, drugdelivery, and/or biological, physiological, and chemical sensing.

The devices disclosed herein can be individually addressable andindependently controlled. Thus, the devices, for example as those usedfor muscular stimulation, can be placed at different locationscorresponding to different muscle groups, and perform stimulation in asynchronized manner. Similarly, brain stimulation devices can be placedat different locations in the brain, and stimulation can be performed ina synchronized manner. The same can be said with drug delivery devices.Moreover, because the devices can be individually addressable andindependently controlled, the devices can be activated and/or poweredasynchronously as well as synchronously. These devices, in certaininstances, can have characteristics dimensions in that the devices aremuch smaller (e.g., one, two, or three orders of magnitude) than theirdepth in tissue. Similarly, the devices, in certain instances, can havecharacteristics dimensions in that the devices are much smaller (e.g.,one, two, or three orders of magnitude) than the source that providesthe power to the devices.

The aspects of the present disclosure, as directed toward apparatuses,devices, and methods, can be utilized alone or in combination withvarious other aspects.

The structures described herein can be used with sensors/devices thatinclude feedback to the sub-wavelength structures. These types ofsensors can include, for example, implantable temperature sensors orimaging devices. In this manner, the devices are responsive to thestructures illustrated above that generate a spatially adaptableelectromagnetic field/signal. The feedback-type devices respond to thepower and/or data portions of the signal provided by the spatiallyadaptable electromagnetic field/signal, and are prompted to respond. Forinstance, temperature sensors located in a patient will broadcast/reportthe temperature of the tissue in response to the power and/or dataportions of the signal provided by the spatially adaptableelectromagnetic field/signal. Additionally, imaging devices implanted ina tissue can broadcast/report the captured images in response to thepower and/or data portions of the signal provided by the spatiallyadaptable electromagnetic field/signal. Moreover, the penetration depthof the spatially adaptable electromagnetic field/signal can be modeledand controlled. Thus, in certain embodiments, the feedback devices canindicate and label data, in response to the spatially adaptableelectromagnetic field/signal, to record the depth at which the device isoperating. By storing this data on a patient-by-patient basis in astorage device, a computer can access and analyze this data forstatistical purposes. By storing the position or label of thefeedback-type device in a memory circuit via a programmable computer,various patient feedback tracking methods can also be realized. Forinstance, the depth of an implantable imaging device can be optimized byanalyzing the surrounding tissue. In this manner, the depth of theimplantable imaging device can be adjusted if it is determined that amore optimal position is possible. Similarly, the depth of animplantable stimulation device can be used to determine the heath of thetissue area surrounding the stimulation device, and determine an optimalpositioning of the device in response to the spatially adaptableelectromagnetic field/signal. Additionally, the feedback-type devicescan respond to the spatially adaptable electromagnetic field/signal andbroadcast data stored in a memory circuit. Thus, the feedback-typedevices can continuously update a physician of the data that is beingtracked by the device. This allows for real-time monitoring, diagnosing,and/or treating a patient wirelessly.

Implantable devices/sensors can be wirelessly powered by controlling andpropagating electromagnetic waves in tissue. The implantable devices canbe implanted in humans or in other animals such as pets, livestock, orlaboratory animals such as mice, rats, and other rodents. Suchimplantable devices/sensors can be implanted at target locations in apatient, as non-limiting examples, to stimulate areas such as the heart,and/or to sense biological, physiological, chemical attributes of theblood, tissue, and other patient aspects. Difficulties in achievingwireless power transfer can occur in the mismatch between the size ofthe implantable devices/sensors and the power transfer source, the depthof the devices/sensors in a patient, and additionally the spatialarrangement of the devices/sensors relative to the power transfersource.

Various aspects of the present disclosure are directed towardapparatuses or methods as exemplified or supported by aspects of theabove noted description/embodiments, as well as thedescription/embodiments of the attached appendices. For instance,certain embodiments of the present disclosure are directed tomanipulation of evanescent fields outside a patient's tissue withsub-wavelength structures to excite/control propagating fields insidethe patient's tissue and thereby generate a spatially focusing andadaptive steering field/signal in the tissue. A sub-wavelength structuregenerates fields that are evanescent in nature near the source. Incontrast, in conventional wireless approaches using inductive coupling,the evanescent components outside tissue (near the source) remainevanescent inside tissue which does not allow for effective depthpenetration.

This disclosure provides embodiments of sub-wavelength structures andmethods for controlling the excitation of those structures to excite thepropagating modes inside tissue from the evanescent modes outsidetissue. As a result, this approach is very effective in transportingenergy to absorption-limited depth inside tissue. The designs disclosedherein include structures that use tissue as a dielectric waveguide totunnel energy into the body. The energy can be received by an implantedmodule which will be discussed below, to allow for wireless powertransfer to implanted devices at depths unattainable with conventionalinductive coupling technology.

This disclosure provides a midfield wireless powering approach thatintegrates an external module configured to transmit wireless power, andone or more implanted modules configured to receive wireless power thatcombines an impulse generator and at least one stimulation electrodetogether into a small, leadless, implantable device. In someembodiments, the implanted module can be small enough to be deliveredvia a catheter or a hypodermic needle. For example, the implanted modulecan be as small as a few millimeters in diameter (2-3 mm) down to havingdiameters on the order of 100's of microns or less. This implantedmodule allows for the transfer of wireless power to nearly any locationin the body at performance levels far exceeding requirements for bothcomplex electronics and physiological stimulation. Because the implantedmodules are small, they can be injected into the targeted nerve ormuscle region directly without the need for leads and extensions, toprovide sensing and stimulation to the targeted nerve, muscle, or tissueregion.

For illustrative purposes, FIGS. 1A-1N show various embodiments andviews of wireless power transmitting modules 100, including one or moresub-wavelength structures 102, consistent with various aspects of thepresent disclosure. A sub-wavelength is defined with respect to thewavelength of the field outside a patient's tissue or in the air. Asub-wavelength structure can be of a dimension less than the wavelengthin air but might be comparable to the wavelength in tissue. For example,at 1.6 GHz, the wavelength in muscle is about 7.3 times smaller than thewavelength in air. Any source structure that is of dimension on theorder of the wavelength in muscle or tissue may be a sub-wavelengthstructure. FIGS. 1C-1E show perspective views of three specificembodiments of wireless power transmitting modules, and FIGS. 1F-1H showside views of those modules, respectively. Similarly, FIGS. 11-IK showperspective views of some wireless power transmitting modules, and FIGS.1L-1N show side views of those modules, respectively.

The sub-wavelength structures of FIGS. 1A-1N can be configured tomanipulate evanescent fields outside a patient's tissue toexcite/control propagating fields inside the patient's tissue togenerate a spatially focusing and adaptive steering field/signal in thetissue. The wireless power transmitting modules 100 shown in FIGS. 1A-1Ncan include the sub-wavelength structure(s) 102 disposed over asubstrate 104 and one or more ground plane(s) 106 (shown in the sideviews of FIGS. 1F-1H and 1L-1N. In some embodiments, the sub-wavelengthstructures 102 can comprise a conductive material, such as a copper. Thesubstrate can comprise an insulating material, such as an epoxy, or aceramic. The substrate can be a solid, rigid, substrate, oralternatively can be a flexible substrate configured to conform to theskin surface of patients. In some embodiments, the sub-wavelengthstructures 100 can further comprise a ground plane bonded to or disposedon the substrate. The ground plane can be disposed on a top surface(FIGS. 1H, 1L, 1N), a bottom surface (FIGS. 1F, 1G), or both top andbottom surfaces (FIG. 1M) of the substrate.

The design of each sub-wavelength structure can be varied depending onthe design requirements of the specific application. FIGS. 1A-1B bothshow a wireless power transmitting module having a plurality ofsub-wavelength structures 102, wherein the sub-wavelength structuresresemble X′ with curved or protruding strips or features. In both theseembodiments, each of the sub-wavelength structures 102 can be excited byone or more independent radio-frequency ports 103 connected to a voltageand/or current source. In some embodiments, the sub-wavelengthstructures can be excited with a voltage ranging from 0.1 V to 10's V,or can be excited with a current ranging from 0.1 A to 10's A. Thefrequency range of the source can range from 300 MHz to 3 GHz. Forappropriate phases between the port signals, the sub-wavelengthstructures can generate circular current paths that mimic the optimalcurrent density. When positioned above tissue, the structures couplepower from the external circuitry into the tissue volume with highefficiency (>90%), as evidenced by both low levels of backside radiationand a pronounced minimum in the scattering parameter spectrum.

Degrees of freedom provided by the phases of the input port signalsenable various interference patters to be synthesized, including thosewith spatially shifted focal regions. Software control of these phasescan refocus the fields without mechanical reconfiguration, which can beuseful for implanted devices inserted on rhythmic organs or forlocomotive devices. In some embodiments, a “greedy” phase searchalgorithm can be implemented based on closed-loop feedback that obtainsfocusing-enhanced power transfer in real-time. In other embodiments, thefeedback signal can be wirelessly transmitted from the implanted deviceto the midfield source.

FIGS. 1C and 1F show a patch sub-wavelength structure 102 c, disposedover a substrate 104 with a ground plane 106 on a bottom surface of thesubstrate. A feed 108 is also shown in FIG. 1F, which is used to feed ortransmit electrical signals to or from the sub-wavelength structure.FIGS. 1D and 1G illustrate a PIFA sub-wavelength structure 102 d,disposed over a substrate 104 with a ground plane 106 on a bottomsurface of the substrate. The feed 108 is shown in FIG. 10, along with ashort 110 connected to the structure 102 d. FIGS. 1E and 1H show a slotsub-wavelength structure 102 e in a ground plane 106 disposed over asubstrate 104. The feed 108 is shown in FIG. 1H. FIGS. 11 and 1L show across slot sub-wavelength structure 102 i in a ground plane 106 disposedover a substrate 104. The feed 108 is shown in FIG. 1L. FIGS. 1J and 1Millustrate an aperture coupled circular slot sub-wavelength structure102 j in a ground plane 106 disposed over a substrate 104. Thisembodiment can further include a ground plane 106 on a bottom surface ofthe substrate. The feed 108 is shown in FIG. 1M. Finally, FIGS. 1K and1N illustrate a half slot sub-wavelength structure 102 k, disposed overa substrate 104 with a ground plane 106 on a top surface of thesubstrate. The feed 108 is shown in FIG. 1N. In all the embodimentsdescribed above and illustrated, one or more power source(s) andamplifier(s) can be connected to the sub-wavelength structure(s) via thefeeds (or ports) to manipulate evanescent fields. Furthermore, in someembodiments, each sub-wavelength structure can include one or more feedsor ports.

The wireless power transmitting modules 100 described above generallyinclude one or more sub-wavelength structures, one or more excitationports, a substrate, and one or more ground planes. The modules 100 canbe controlled by a controller (both hardware and software) todynamically shift a focal region of the electromagnetic field.

Some discussion on various techniques for transferring wireless powerwill now be described. FIG. 2 shows the magnetic field 212 generated bya conventional inductively coupled loop source 214, in both the yz andxz planes. As can be seen, the magnetic field is generated perpendicularto the tissue interface 216, and is parallel with the direction ofdesired wireless power transfer to an implant disposed in tissue belowthe loop source, such as an implanted device 218.

FIG. 8 shows several arrangements of sub-wavelength structures andrepresentative tissue, consistent with various aspects of the presentdisclosure.

In contrast, FIGS. 3A-3E show the magnetic fields 312 produced byvarious sub-wavelength structures of the present disclosure. Thesestructures generate a magnetic field 312 parallel to the tissueinterface 316, and perpendicular to a propagating wave generated intissue that transmits wireless power to an implanted device 318. FIG. 3Ashows the magnetic field generated with a patch sub-wavelength structure302 c (FIGS. 1C and 1F) in the yz and xz planes. FIG. 3B shows themagnetic field generated with a PIFA sub-wavelength structure 302 d(FIGS. 1D and 10) in the yz and xz planes. FIG. 3C shows the magneticfield generated with a cross slot sub-wavelength structure 302 i (FIGS.11 and 1L) in the yz and xz planes. FIG. 3D shows the magnetic fieldgenerated with an aperture coupled circular slot structure 302 j (FIGS.1J and 1M) in the yz and xz planes. FIG. 3E shows the magnetic fieldgenerated with a half slot sub-wavelength structure 302 k (FIGS. 1K and1N) in the yz and xz planes.

FIG. 4A shows a wireless power transmitting system including a wirelesspower transmitting module 400 and an implanted device 418 inside a humanbody. In FIG. 4A, the device is shown implanted in a chest cavity of thepatient, such as in or near the heart. It should be understood from thisfigure that the implanted device can be placed anywhere in the body,such as in the heart, brain, lungs, spinal cord, bones, nerves, sinuses,nasal cavity, mouth, ears, peritoneal cavity, arms, legs, stomach,intestines, digestive tract, kidneys, bladder, urinary tract, or anyother organ or part of the body that can benefit from the sensing and/orstimulation features provided by the systems described herein.

In FIG. 4A, the transmitting module 400 can be positioned above the skinof the patient, and the implanted module comprising a receive coil canbe implanted in the patient. Power transfer occurs when the interactionof the source fields with the coil structure results in work extruded bya load in the implanted module. For a sub-wavelength coil, only thelowest order mode is important and the transfer mechanism can bedescribed by electromagnetic induction characteristics of dynamicmagnetic field interactions. The electric and magnetic fields generatedby a time-harmonic current density Js on the surface of the sourceconductor can be solved by decomposing the current density into itsspatial frequency components. Each component corresponds to a plane wavewith propagation determined by phase matching conditions for refractionand reflection over planar boundaries, from which the total field intissue can be recovered at each depth z by integration over the sourcespectrum.

The properties of the mid-field region are key to optimal powering. Thesub-wavelength structures manipulate evanescent fields to excite/controlpropagating waves (alternating electric and magnetic fields) and therebygenerate a spatially focusing and adaptive steering field/signal intissue that converges on the implanted device. Back-propagation offields at the focal plane to the surface of the skin reveals that thesource is highly oscillatory and composed of significant evanescentcomponents that are important only in the near-field. In contrast withconventional near-field powering, however, these evanescent componentsexcite propagating modes in tissue that transport energy toabsorption-limited depths.

FIGS. 4B-4C show the difference between the ability of a near-field orinductively coupled wireless power transfer system (FIG. 4B) to transferpower into a depth of tissue compared to the mid-field design (FIG. 4C).of the present disclosure. As seen in FIG. 4C, the mid-field design ofthe present disclosure allows for transmission of wireless power to adepth in tissue not attainable by inductively coupled systems.

In some embodiments, a focal point of the wireless power transfer systemof the present disclosure can be adjusted to change a direction of thepropagating wave. FIG. 4C illustrates formation of a propagating wave ina direction directly below the external module, along line 419 a.However, in some embodiments, the focal point can be adjusted to causethe propagating wave to travel in a steer direction through the tissue,such as along lines 419 b or 419 c. This adjustment can be attained byadjusting a phase and/or amplitude of one or more of the sub-wavelengthstructures of the external module.

FIGS. 5A-5B shows two embodiments of architectures for a controller ofthe wireless power transmitting modules described herein, for excitingthe ports of the sub-wavelength structures. These architectures can beconfigured to control one or more sub-wavelength structures 502 a-502 nof the wireless power transmitting modules. In each architecture, the RFsignal can be sourced from an oscillator 520, and be dividedsymmetrically into multiple RF signals through a power divider. In thearchitecture of FIG. 5A, the signal is then fed through attenuator(s)522 with variable controllable attenuation settings. The signals canthen be fed through phase shifter(s) 524 with controllable phase, andthen amplified with amplifier(s) 526. This architecture producescontrolled phase and amplitude signals at each port of the module. Thearchitecture on in FIG. 5B is configured to produce the same controlledphase and amplitude signals, but with fewer components by combining theamplifier(s) and the amplitude control element(s) into a singlecomponent 528.

Implanted Module.

One embodiment of an implanted module for receiving wireless power isshown in FIG. 6. The implanted module can include a coil 630 disposedover an integrated chipset (IC) 632. The coil 630 can be a loop (ormultiple loops) of a conductor. In some embodiments, the coil 630 has adiameter of less than 2 mm. The coil can be configured to receive thewireless power transmitted from the external modules described herein.The module can optionally include features 634 for sensing and/orstimulating tissue, such as electrode(s) or sensors. The electrodes cancomprise, for example, screw-type electrodes, planar electrodes, or cuffelectrodes. In other embodiments, the sensors can comprise biopotentialsensors, pressure sensors, 02 sensors, etc. The implanted module canoptionally include electrical components for the storage of energy, suchas a capacitor or battery 636. Due to the small size of the implantedmodule (2 mm or less in diameter), the implanted module can be deliveredand implanted into a patient with minimally invasive techniques, such aswith a catheter 638, a cannula, a needle, or the like.

Because the power levels supported by a midfield wireless poweringapproach far exceed requirements for microelectronic technologies (e.g.,in one embodiment, an input power level of 500 mW from the externalmodule can deliver approximately 200 uW of power over 5 cm of tissue toa 2 mm diameter implant coil), more sophisticated functions can beimplemented such as real-time monitoring of chronic disease states orclosed-loop biological sensing and control by the implanted module.Hence, in some embodiments, the implanted module can include one or moreof the following building blocks:

Power management. To increase the efficiency of rectification and powermanagement of wirelessly powered implants operating in theelectromagnetically weakly coupled regime, AC-DC conversion circuits inthe implanted module can be divided into the low-voltage andhigh-voltage domains. FIG. 7A shows an architecture that can be includedin the IC of the implanted module to handle the power managementfeatures of the implant. FIG. 7A shows a coil 730 electrically connectedto one or more capacitors (or variable capacitors) 740, multistagerectifiers 742, and regulators 744, to divide the AC-DC conversioncircuits into low-voltage and high-voltage domains.

Battery Storage.

A rechargeable battery such as thin film battery can be included in theimplanted module for temporary energy storage and for use as anefficient charge pump for the power management circuitry. In someembodiments, the thin film battery can be stacked to increase the energydensity.

Power Detection.

The instantaneous power level received by the implanted module can bedetected and sent via a data transmitter to the external module foradaptive focusing onto the implant module in the midfield. Data can betransmitted between the implanted module and the external module througha wireless link. In some embodiments, the wireless link can operate inthe frequency range of the power transmission, or in other embodiments,the wireless link can operate in a different frequency range. Thedetected power level can be used in a closed-loop feedback control withthe controller of the system to adjust and focus the external module foroptimal wireless power transfer.

Pulsed RF Modulation.

Conventional load modulation does not work in the midfield due to thelow quality factor of the implant antenna, leading to poorsignal-to-noise ratio and substantial link margin fluctuation. Toovercome this problem, the data transmitter of the implanted module canuse pulsed RF modulation. To ease detection at the external module, thedata and power carriers can operate at different center frequencies.

Programmable Current Drivers.

Stimulation applications differ mainly by the characteristics of theelectrical pulses such as intensity, duration, frequency, and shape. Thecurrent drivers for stimulation are designed to support wide range ofthese parameters and can be programmed via the wireless data link. Thecurrent drivers can also be configured to support actuation such aslocomotion.

Programmable Digital Core.

The digital core coordinates the interaction among various blocks in theimplanted module, communication between the implant and externalmodules, and the multi-access protocols. Each implant module can haveits own identification (ID) such as via an ID stored in the memory ofthe implanted module.

Data Receiver and Transmitter.

The external module can remotely communicate with each implanted moduleto program or configure each implanted module via the data receiver.FIG. 7B shows one embodiment of a data receiver based on envelopdetection and a data transmitter based on an ultra-widebandarchitecture. The receiver and transmitter can be time multiplexed by aT/R switch 746 connecting to the power receiving coil or to a separateantenna. Each implanted module can have its own ID 748 for multi-access.A digital controller 750 can be implemented to handle the multi-accessprotocol 752, commands from the external module, and feedback data tothe external module.

Sensing Frontend.

The sensing frontend can comprise pre-amplifiers, analog-to-digitalconverters (ADC) to discretize signals from the pre-amplifiers, anddrivers for the sensors. Signals from the output of the ADCs can eitherbe stored in the non-volatile memory of the implanted module or sent tothe external module via the Pulsed RF modulator. In addition, the sensedsignals can provide biological feedback for adjusting parameters of thecurrent drivers. FIG. 7C shows the architecture for one or multiple LEDdrivers, and the electrical sensing and stimulation frontends. The LEDdrivers can be connected to LEDs for optical stimulation of tissue(nerves). The electrical sensing and stimulation frontends can also beconnected to electrodes for sensing the biological activities andaltering the electrical pathways.

Non-Volatile Memory.

Flash memory, for example, can be included to record usage model of theimplant module such as the time of activation and setting of the currentderiver, and/or to store measurements from the sensing frontend.

Modular Construction.

The implanted module can be customizable depending on the particularneeds or requirements of the end user. For example, the implanted modulecan include a number of base components including the wireless powerreceiving coil and the IC, and can further include an interface that canreceive any type of sensor or stimulator desired by the user. Forexample, the implanted module can be configured to receive any type ofsensor, such as thermal, chemical, pressure, oxygen, PH, flow,electrical, strain, magnetic, light, or image sensors, or any type ofstimulator, such as electrical, optical, chemical, or mechanicalstimulators, or a drug delivery apparatus. The modular approach of theimplanted module can therefore be customized to accommodate theparticular needs of the user.

All the above building blocks in the implanted module can be integratedinto a single die as system-on-chip (SoC) or multiple dies enclosed in asingle module as system-in-package (SiP).

External Module.

The external module (described above) can be configured to energize andcontrol the implanted modules, and to perform noninvasive readoutthrough a bidirectional wireless link setup with the implanted modules.The external module can include one or more of the following buildingblocks:

Midfield Coupler.

FIGS. 1A-1N show various shapes and patterns for the external module ormidfield coupler, which can include one or more sub-wavelengthstructures. The coupler can be made on solid substrate, or on a flexiblesubstrate configured to conform to the skin surface of patients.

Dynamic Midfield Focusing Circuits and Algorithms.

Based on the power measurement feedback from the implant module, theexternal module can run an algorithm, for example, the greedy searchalgorithm, to change the phase and/or magnitude settings in each elementof the midfield coupler so as to dynamically shift the focal region tothe individual implant module. For example, the implanted module candetect a power level of received wireless energy, and the externalmodule can automatically adjust the phase and/or amplitude of thesub-wavelength structures to adjust the focal point of the transmittedenergy signal. This adjustment can be made automatically and in realtime to optimize wireless power transmission between the external moduleto the internal module.

Bidirectional Wireless Link to the Implant Module.

The wireless link can activate the implanted module, program the settingof the implanted module, and download measurements from the sensingfrontend of the implanted module. The data rate for the downlink; fromthe external module to the implanted module, can be a few Mbps or lower,while the data rate for the uplink; from the implant module to theexternal module should be higher, can be in the range of 1 Mbps or evenhigher.

Multiaccess Protocols.

These protocols can coordinate the implanted modules to carry outsynchronous tasks such as coordinated multi-site stimulation. In someembodiments, multi-access schemes can be time multiplexing and frequencymultiplexing.

Patient/Clinician User Interface.

A peripheral device including a display can be integrated with theexternal module to interface with a patient and/or clinician. In otherembodiments, the integrated peripheral device can be replaced by abidirectional wireless link communicating with a smartphone or a tablet.In this embodiment, the patient and clinician can interface with theexternal module using the display of the smartphone or tablet throughthe wireless link.

In some embodiments, the entire external module can be integrated into apalm-size device and held by the patient for on-demand applications. Itcan also be worn on the body or affixed to the skin surface. Patientscan use the external module to charge the battery of the implant modulesas needed. In some embodiments, the implanted module(s) can be chargedwith only a few minutes of wireless charging per week/month. Duringcharging, patients can also download usage record from the implantmodules and send the record to the clinician for analyses.

Various aspects of the present disclosure are directed toward poweringof multiple active implantable sensors or devices using a single powersource. The types of implantable devices/sensors that can be poweredusing the single power source, consistent with various aspects of thepresent disclosure, are numerous. For instance, the implantable devicescan be used for muscular stimulation, stimulation/sensing to regulate apatient's heart beat, multisite deep brain stimulation, drug delivery,and/or biological, physiological, and chemical sensing. The systemsdescribed herein can also be configured to be used in the followingapplications:

Cardio Pacemaker.

The implanted module can be delivered via a catheter through thevasculature into the right ventricle of a patient. A separate implantedmodule can be delivered through the coronary sinus into the coronaryvein, and placed on the left ventricular epicardium. These implantedmodules can include stimulation and sensing electrodes to apply leadlesspacing to the heart. Thus, leadless biventricular pacing can be achievedwith the present system with only minimally invasive procedures. Inaddition, the procedure time can be shortened substantially over priorapproaches. This can also eliminate any complication during to themultiple leads and extensions.

Deep-brain stimulation. Current procedure involves the drilling of holeswith diameter >1 cm in the skull to insert a lead and the extension fromthe lead to the stimulating module. Due to the invasiveness of theprocedure, only a limited number of target sites are selected forplacing the electrodes. By contrast, the implanted modules in thisdisclosure, being very small, can be injected into the brain via otherless invasive routes. Since there is no lead and extension wire in thepresent system, more target sites for stimulation can be supported. Thisresults in less infection and lower regulatory risk.

Spinal cord stimulation. Batteries in newer models of spinal cordstimulator are rechargeable due to the high power requirement. However,their powering approaches are exclusively based on inductive coupling(or near-field coupling). Since the harvesting components are large inthese systems, they can only be placed subcutaneously. Therefore, thelead and extension wires in these systems potentially restrict thelocation of the electrodes for effective stimulation. In thisdisclosure, the power-harvesting component in the implanted module isrelatively tiny. The entire implanted module can be easily placed nextto the targeted nerve region in the spinal cord and requires no leadwire connecting them. This results in less infection, less damage to thespinal cord tissue, and more effective stimulation.

Peripheral nerve stimulation. Most current devices support low-frequencystimulation and only a few of them support high-frequency low-intensitystimulation due to the much higher power requirement. The systems ofthis disclosure can support both modes. In addition, the bidirectionalwireless link provides instant programmability, switching betweendifferent modes.

Stimulation to treat obstructive sleep apnea (OSA). The implantedmodules of this disclosure can be injected and directly embedded intothe muscular tissue near the tongue, and can deliver electricalstimulation to open the airway of a patient during sleep. Multipleimplant modules can be injected into different muscular groups tointensify the muscle contraction. When needed, patients can charge theimplanted modules with the external module and simultaneously, downloada time stamp of each OSA episode. This information can be sent to theclinicians. Data collected can also be used to reprogram the implantedmodules.

Medical sensors. Batteryless implanted sensors are typically passive innature, that is, there is no active circuitry in the device to conditionthe sensed signals. To compensate for the poor signal quality, anexternal reader is needed to be very sophisticated and is usually large(cannot be fitted on a palm). In addition, not many stimuli can bedetected by passive sensors. The lack of active implanted sensors ismainly due to the lack of an efficient wireless powering approach. Forexample, the inductive coupling approach used in the rechargeableimpulse generator for spinal cord stimulation has limited penetrationand the receiver (the implanted device) is large. The system of thepresent disclosure allows for the transfer of substantial amount ofpower to small implanted modules at nearly any location in the body froma palm-size external module. This enables an array of new sensingapplications for continuous monitoring in the medical field, forexample, post-surgery oxygen sensing in the heart and the brain.

Wireless endoscopes. Current capsule endoscope has limited batterylifetime, leading to incomplete small-bowel examination which is one ofthe major clinical failures. The implant module in our invention issmall and has indefinite power supply, solving the deficiency of currentendoscopes. In addition, since our implant module is many times smallerthan the current capsule endoscope, patients can swallow multiple of theimplant modules simultaneously. They are expected to orient differentlyin the intestine and therefore, can take pictures from different anglesat the same location, improving the field of view. The images collectedfrom them will improve the diagnosis. Finally, the probability ofretention is expected to be dramatically reduced, avoiding the need ofsurgical or endoscopic retrieval.

Implanted drug delivery. Current implanted drug delivery systems arelarge and mostly cannot be placed local to the site that the drug isneeded. Based on this disclosure, the implanted module can be injectedinto a targeted tissue region (for example, a tumor) where the drug isneeded. The implanted module can include a number of drug reservoirs.The drug reservoirs can be activated by the external module via thepatient/clinician user interface to release a drug into the targetedtissue region.

Temporary treatment. Currently, screening tests are typically performedbefore a permanent impulse generator is implanted. During the screeningtest, a patient may receive a temporary, external impulse generator. Thegenerator can connect to an extension and a lead that are surgicallyplaced in the body. In this period, the external impulse generatorcollects patient usage data and efficacy of the treatment. However,according to this disclosure, the implanted module having an electrodeand an impulse generator can be injected into the targeted nerve/muscleregion, eliminating the need for a temporary generator with leads. Thereis therefore no need for the external temporary impulse generator. Inaddition, this disclosure can also replace the temporary sensing andpacing leads used in patients after cardiac surgery.

Laboratory Experiments. The implanted module can be injected into labanimals or rodents (such as mice, rats, etc.) to monitor or senseparameters of the animal and/or provide stimulation to the animal in anexperimental setting. The small size of the implanted module canadvantageously provide opportunities to monitor the animal that has notbeen previously available. For example, the implanted module could beimplanted on or near the brain of a rodent to monitor electrical signalsof the brain. The implant can be wirelessly powered with the externalmodule described above, and can be configured to communicate informationback to the external module relating to the animal.

The devices are individually addressable and independently controlled.Thus, the devices, for example as those used for muscular stimulation,can be placed at different locations corresponding to different musclegroups, and perform stimulation in a synchronized manner.

Similarly, brain stimulation devices can be placed at differentlocations in the brain, and stimulation can be performed in asynchronized manner. The same can be said with drug delivery devices.Moreover, because the devices can be individually addressable andindependently controlled, the devices can be activated and/or poweredasynchronously as well as synchronously. These devices, in certaininstances, can have characteristics dimensions in that the devices aremuch smaller (e.g., one, two, or three orders of magnitude) than theirdepth in tissue. Similarly, the devices, in certain instances, can havecharacteristics dimensions in that the devices are much smaller (e.g.,one, two, or three orders of magnitude) than the source that providesthe power to the devices.

The aspects of the present disclosure, as directed toward apparatuses,devices, and methods, can be utilized alone or in combination withvarious other aspects.

For information regarding details of other embodiments, experiments andapplications that can be combined in varying degrees with the teachingsherein, reference may be made to the experimental teachings andunderlying references provided in the following attachments which form apart of this patent document and are fully incorporated herein byreference. Embodiments discussed in these appendices are not intended,in any way, to be limiting to the overall technical disclosure, or toany part of the claimed disclosure unless specifically noted.

In such contexts, these building blocks and/or modules representcircuits that carry out one or more of these or other relatedoperations/activities. For example, in certain embodiments discussedabove, one or more blocks and/or modules are discrete logic circuits orprogrammable logic circuits configured and arranged for implementingthese operations/activities, as in the circuit modules/blocks describedabove and in the Appendices. In certain embodiments, the programmablecircuit is one or more computer circuits programmed to execute a set (orsets) of instructions (and/or configuration data). The instructions(and/or configuration data) can be in the form of firmware or softwarestored in, and accessible from, a memory (circuit).

In connection with the above discussed features and illustrativefigures, such structures can be used with sensors/devices that includefeedback to the sub-wavelength structures. These types of sensors caninclude, for example, implantable temperature sensors or imagingdevices.

In this manner, the devices are responsive to the structures illustratedabove that generate a spatially adaptable electromagnetic field/signal.The feedback-type devices respond to the power and/or data portions ofthe signal provided by the spatially adaptable electromagneticfield/signal, and are prompted to respond. For instance, temperaturesensors located in a patient will broadcast/report the temperature ofthe tissue in response to the power and/or data portions of the signalprovided by the spatially adaptable electromagnetic field/signal.Additionally, imaging devices implanted in a tissue can broadcast/reportthe captured images in response to the power and/or data portions of thesignal provided by the spatially adaptable electromagnetic field/signal.Moreover, the penetration depth of the spatially adaptableelectromagnetic field/signal can be modeled and controlled. Thus, incertain embodiments, the feedback devices can indicate and label data,in response to the spatially adaptable electromagnetic field/signal, torecord the depth at which the device is operating. By storing this dataon a patient-by-patient basis in a storage device, a computer can accessand analyze this data for statistical purposes.

By storing the position or label of the feedback-type device in a memorycircuit via a programmable computer, various patient feedback trackingmethods can also be realized. For instance, the depth of an implantableimaging device can be optimized by analyzing the surrounding tissue. Inthis manner, the depth of the implantable imaging device can be adjustedif it is determined that a more optimal position is possible. Similarly,the depth of an implantable stimulation device can be used to determinethe heath of the tissue area surrounding the stimulation device, anddetermine an optimal positioning of the device in response to thespatially adaptable electromagnetic field/signal. Additionally, thefeedback-type devices can respond to the spatially adaptableelectromagnetic field/signal and broadcast data stored in a memorycircuit. Thus, the feedback-type devices can continuously update aphysician of the data that is being tracked by the device. This allowsfor real-time monitoring, diagnosing, and/or treating a patientwirelessly.

Wireless Powering for Catheter-Insertable Electronics

Seamless integration of electronics into the body can restore andaugment many physiological functions. However, its realization isrestricted by the enormous mismatch in scale between the microscopic(nanometers to millimeters) integrated electronics and macroscopic(centimeters) energy storage or harvesting components. Here, weintroduce midfield wireless powering, which we validate to be capable ofpowering a 2-mm active device 10 cm deep in tissue three orders ofmagnitude reduction in size with a tenfold increase in depth overconventional approaches. The powering source is realized through apatterned metal plate that excites propagating waves in tissue from itsevanescent components, and generates a focused and spatially adaptiveelectromagnetic midfield. To illustrate the capabilities of thisapproach, we built a 2-mm diameter, 70-mg wireless pacemaker anddemonstrated closed-chest wireless pacing.

We report power transfer to miniaturized semiconductor devices byexploiting propagating waves in biological tissue generated in themidfield electromagnetic region of the source. At the scale of amillimeter, wireless devices must operate at depths in dissipativetissue that are over an order of magnitude greater than theircharacteristic sizes. In such configurations, established mechanisms forfree-space transfer are highly inefficient: near-field approaches, whichrely on strong coupling occurring between objects with matchedelectrical characteristics such as resonances and impedances, do notgeneralize easily to geometries with extreme size asymmetry, whilefar-field transfer is limited by absorption over surfaces of the body.Although energy sources based on thermoelectric, piezoelectric,biopotential, or glucose harvesting are promising alternatives, they donot in their existing forms (<0.1 μW/mm²) achieve power densitiessufficient for a millimeter-sized device. Powering in the midfieldprovides a different approach in which both evanescent and radiativecomponents of a structure are coupled to modes in tissue that transportenergy continuously away from the source. Interference resulting fromphase differences between these components affords additionalopportunity for spatially focused and dynamically adjustable fieldpatterns inside tissue.

We demonstrate systems that exploit these characteristics to wirelesslypower miniaturized devices—sufficiently small to be delivered by acatheter—inserted several centimeters in heterogeneous tissue. For manyclasses of electronics, the level of performance obtained by thisapproach exceeds requirements for advanced monitoring and controlcapabilities that might be developed for applications in medicine,neuroscience, or human-machine interfaces.

We identify the physics underlying the midfield by considering powertransfer through a multilayer approximation of the chest wall. Thepowering configuration consists of a source positioned above the skinand a receive coil inserted in the cardiac tissue layer (FIG. 1A). Thelayered structure permits a simple description of wave propagation whileproviding some physical resemblance to surfaces on the body. Powertransfer occurs when the interaction of the source fields with the coilstructure results in work extracted by a load in the receiver circuit.For a subwavelength coil, only the lowest order mode is important andthe transfer mechanism can be described by electromagnetic inductioncharacteristic of dynamic magnetic field interactions. In this case, thepower transferred from the source to the coil is given by

$\begin{matrix}{P_{SC} = {\int{d^{3}{{{rM}_{C}(t)} \cdot \ \frac{d\; {B_{s}(t)}}{dt}}}}} & (1)\end{matrix}$

where Bs is the magnetic field generated by the source and Mc theinduced magnetization due to current in the coil.The electric and magnetic fields generated by a time-harmonic currentdensity Js on sur-face of the source conductor can be solved bydecomposing the current density into its spatial frequency components(kx, ky). Each component corresponds to a plane wave with propagationdetermined by phase matching conditions for refraction and reflectionover planar boundaries, from which the total field in tissue can berecovered at each depth z by an integration over the source spectrum.Using phasor notation with a time dependence of exp(−iwt), we can definean efficiency in terms of these fields

$\begin{matrix}{\eta = \frac{{{\int{d^{3}{{rM}_{C}^{*} \cdot B_{S}}}}}^{2}}{\left\lbrack {{\int{d^{3}r\mspace{11mu} {Im}}} \in {(\omega){E_{S}}^{2}}} \right\rbrack \left\lbrack {{\int{d^{3}r\mspace{11mu} {Im}}} \in {(\omega){E_{C}}^{2}}} \right\rbrack}} & (2)\end{matrix}$

Formally, η is the ratio of power available at the coil to the totalabsorbed power. Eq. 2 considers only dissipation in tissue: otherlosses, such as radiative and ohmic loss, arise in practice, but theamount of power that can be coupled into the body is essentially limitedby electric field-induced heating in tissue. Efficiency as defined aboveis intrinsic to the fields in the tissue half-space and gives an upperbound on the efficiency that can be obtained. The work extracted by theload can be determined by impedance matching considerations and otherimplementation dependent factors.

The choice of source Js that maximizes efficiency in Eq. 2 is key forefficient power transfer. Direct search over the space of such candidatecurrent densities, however, is not feasible because it leads tointractable computational complexities. Approaches to related problems,such as the design of optical antennas, rely on non-global optimizationstrategies (e.g. genetic algorithms) that find locally optimalstructures. We developed an alternative method that enables the globaloptimum to be analytically solved, although in terms of a non-physicalcurrent density, for a specified powering configuration. In thisapproach, an electric current density is defined with componentstangential to a plane between the source structure and the tissue. Forevery source, the electromagnetic equivalence theorem enables such atwo-dimensional current density to be chosen from the overall set S thatis indistinguishable in the lower z<0 half-space from the physicalsource of the fields. Remarkably, a solution to the optimization problemmaximize JsεS η (Js) can be found in closed-form as a consequence of thesimple vector space structure of S. Unlike solutions derived from localoptimization algorithms, the maximum efficiency obtained by thisprocedure is a rigorous bound on the performance that can be achieved byany physical realization of the wireless powering source. By exploringsuch global solutions across a range of frequencies with appropriatedispersion models for biological materials, we were able to establishoptimal power transfer for the chest wall structure. For a 2-mm diametercoil inserted at a 4 cm depth in tissue with its magnetic dipole momentaligned with the x axis to maximize coupling with a transverse magneticfield component, the optimal frequency occurs near 1.6 GHz where thetissue wavelength (λ≈2 to 4 cm) is comparable to the powering distance(FIG. 12). The properties of this midfield region are key to optimalpowering, as the resultant magnetic field in FIG. 9A shows. The fieldsconsists of propagating waves (alternating electric and magnetic fields)that converge on the device. Back-propagation of fields at the focalplane to the surface

of the skin reveals that the source is highly oscillatory and composedof significant evanescent components √{square root over (k_(x) ²+k_(y)²)}>k₀ that are important only in the near-field (FIG. 9B). In contrastwith conventional near-field powering (FIG. 1C), however, theseevanescent components excite propagating modes in tissue that transportenergy to absorption-limited depths. Although far-field methods alsoexploit such propagating waves, they only involve √{square root over(k_(x) ²+k_(y) ²)}>k₀ components that do not provide similar spatialconfinement of electromagnetic energy (FIG. 1D). The amplitudes andphases of the components emerging from the optimization solution aredesigned such that their superposition is focused at a device plane. Therequisite near-field characteristics (FIG. 1E) are unlike those ofdipole or coil primitive elements, requiring more complexelectromagnetic structures for synthesis.

FIGS. 9A-9E show a multilayer model of power transfer to a subwavelengthcoil in tissue. FIG. 9A shows a schematic of midfield power transfer ina multilayer tissue structure. The magnetic field (Hx, linear scale)corresponding to the analytically derived optimal source current densityJs (f=1.6 GHz, Z_(skin)=−1 cm, Z_(coil)=−5 cm) is shown; the receivecoil is oriented in the x direction. FIG. 9B shows spatial frequencyspectra at specified depth planes. k₀ is the wavenumber corresponding topropagation in air; k_(muscle) the wavenumber in muscle tissue. FIG. 9Cshows a magnetic field generated by conventional near-field powertransfer (Hz, logarithmic scale, f=10 MHz). FIG. 9D shows a magneticfield generated by far-field power transfer calculated by removing theevanescent components in FIG. 9B (Hz, logarithmic scale). FIG. 9E showsa cross-section view of FIG. 9A (Hz, logarithmic scale). FIGS. 9C to 9Eare normalized such that the maximum electric field in tissue is thesame.

Near-perfect efficiency, while possible for mid-range systems with largeand symmetric coils, is not necessary for realizing wirelesscapabilities. Many classes of low-power integrated circuits incorporatefunctions such as communication, sensing, and stimulation in devicesconsuming less than 20 μW during active operation. In the presentnumerical example, the theory indicates that up to 320 μW can betransferred to the load if the source couples 500 mW, roughly the powerradiated by a cell phone, into the chest (η=6.4×10⁻⁴). The efficiency ofpower transfer is one to two orders of magnitude greater than that ofnear-field systems in similar configurations, which are less than 10⁻⁵due to weak coupling. The latter performance is generally not sufficientfor operation in close proximity to tissue—safety and complexityconsiderations limit the power flowing through the external sourcecircuitry (<2-10 W depending on frequency)—but such requirements can bemet with the above theoretical transfer characteristics.

FIGS. 10A-10G show midfield power transfer realization with a patternedmetal plate. FIG. 10A shows a schematic of the source design (dimensions6 cm×6 cm, operating frequency 1.6 GHz) and the magnetic field H_(x)over the skin surface. FIG. 10B shows field patterns with spatiallyshifted focal points designed by adjusting relative phases between theport signals. The upper diagrams in FIG. 10B show formation of thepropagating wave in a direction directly below the external module. Thelower diagrams in FIG. 10B show adjustment of the focal point of thewave, and thus the direction of the wave. As described, this adjustmentcan be attained by adjusting a phase and/or amplitude of thesub-wavelength structures of the external module.

FIG. 10C shows spatial frequency spectrum along the k_(x) axis for themagnetic field in FIG. 10A compared with the theoretical optimum. FIG.10D shows an experimental setup for power transfer measurements. FIG.10E shows theoretical, numerically simulated, and measured powerreceived by a 2-mm coil when coupling 500 mW into a liquid solutionmimicking the dielectric properties of muscle tissue. FIG. 10F shows astrobed position of the LED as the wireless device moves in a “S” shapedtrajectory. A real-time control algorithm enables dynamic focusing whilethe device is in motion. FIG. 10G is the same as FIG. 10F but withoutdynamic focusing; the field pattern is static and focused at the center.

Our physical realization of the midfield powering source consists of ametal plate patterned with slot structures and excited by fourindependent radio-frequency ports (FIG. 10A). For appropriate phasesbetween the port signals, the slot-array structure generates circularcurrent paths that mimic the optimal current density. When positionedabove tissue, the structure couples power from the external circuitryinto the tissue volume with high efficiency (>90%), as evidenced by bothlow levels of backside radiation (FIG. 10B) and a pronounced minimum inthe scattering parameter spectrum (FIG. 11).

FIG. 11 shows a reflected power spectrum S₁₁ as a function of frequencymeasured at an input port of the slot-array source. When positioned 1 cmabove the human chest or porcine tissue, a pronounced dip is observed atthe design frequency 1.6 GHz. Since backside radiation is negligible,electrical power from the signal generator is efficiently coupled topropagating waves in tissue. At a higher operating frequency 1.9 GHz,power is efficiently coupled into waves in air. The structure behaves asa conventional radiative antenna in this region.

Analysis of the fields on the surface of the skin shows that theevanescent spectrum approximates the theoretical optimum (FIG. 10C),although the contribution of the radiative modes are about a factor oftwo greater owing to the inherent directionality of the planarstructure. Regardless, when transferring power to a device submerged ina 0.5% saline solution with dielectric properties mimicking muscletissue, experimental and numerical studies show that the design obtainsefficiencies within 10% of the theoretical bound (FIG. 10E).

Degrees of freedom provided by the phases of the input port signalsenable various interference patterns to be synthesized, including thosewith spatially shifted focal regions (FIG. 10B). Software control ofthese phases can refocus the fields without mechanical reconfiguration,which could be useful for probes inserted on rhythmic organs or forlocomotive devices. We implemented a “greedy” phase search algorithmbased on the closed-loop feedback relayed over an fiber optic cable(FIG. 10D) that obtains focusing-enhanced power transfer in real-time.Over a “S” shaped trajectory of motion, this adaptation eliminates theoutage regions that occur in the static case (FIG. 13), indicating acoverage area much wider than that intrinsic to the focal region (FIGS.10F and 10G). Incorporating components for wireless communication in thedevice will enable an untethered realization of this and other relatedcontrol algorithms.

Received power and array phases as the device moves along the “S”-shapedtrajectory. FIG. 13A shows a trajectory of the device in a liquidsolution with dielectric properties imitating muscle tissue. The dotmarks the starting position. FIG. 13B shows separation between thedevice (dot) and the center of the source (white dot). The distance isapproximately 6 cm, including a 1 cm air gap between the source and theliquid. FIG. 13C shows power received by the device measured by theflashing rate of the LED. The minimum power to operate the device isabout 10 pW. The dynamic phase adaption algorithm enables higher levelsof power to be transferred as the device moves. FIG. 13D shows a phaseof each port, relative to a phase stationary port 4, controlled by thealgorithm along the trajectory of motion.

Relative to a millimeter-sized coil, probes based on conductive wiresinteract strongly with the source fields, making them unsuitable formeasuring the transferred power. We instead developed an integratedprobe that separates the powering and measurement modalities by encodingthe electrical power level into the frequency of optical pulses. Theprobe consists of a multi-turn coil structure, rectifying circuits forAC-DC power conversion, a control unit for regulating the pulseamplitudes, and a light-emitting diode (LED) (FIGS. 14A and 14B). Afiber optic cable guides the signal to a terminating photodiode at themeasurement location. We are able to calculate the power transferred tothe coil, calibrated for circuit-dependent inefficiencies, by recordingthe end-to-end power levels at multiple reference flash rates. From thenon-linear response of the circuit, a system of linear equations can besolved to yield the wireless power transfer efficiency in Eq. 1.

We evaluate the performance of our powering scheme in complex tissuegeometries by designing two configurations that simulate power transferto devices in the left ventricle of the heart and the cortex region ofthe brain in a pig. The source and device positions within the tissuevolume using magnetic resonance imaging (MRI) reconstructions, areseparated by at least 4 cm of heterogeneous tissue. When coupling 500 mWinto tissue, we estimate the power transferred to the coil to be 195 μWfor the heart and 200 μW for the brain configurations. The overallefficiency, inclusive of all subsequent circuit losses, is about afactor of two less, depending on environmental loading effects: wemeasured the respective powers flowing through the pulse control unitsto be 90 μW and 82.5 μW. Most of the intermediate losses are due toinefficiencies in power conversion and can be expected to besubstantially reduced with improved rectifier designs.

To determine the spatial distribution of absorbed energy, we operatedthe source over an extracted piece of tissue and measured surfacechanges in temperature. Infrared imaging shows about a maximum 1.7° C.increase when 500 mW is continuously coupled into tissue over a periodof 1 hour. The heating in live subjects is expected to be substantiallyless because of additional mechanisms for thermal regulation. From theinitial rate of change in temperature, we estimate that the peakspecific absorption rate (SAR) is 3.9 W/kg in the volume directly underthe source. As a cross-check, this value can be compared to the peak SARthat occurs when the same amount of power is coupled into a multilayertissue structure. We numerically calculate the SAR to be 3.5 W/kg, whichis within 15% of the experimental estimate. Although the measuredabsorption exceeds that of cell phones, it is approximately a third ofthe threshold for controlled environments and remains well below otherlevels specified in medical guidelines, such as those for MRI. If thepower coupled into tissue is allowed to meet the maximum permitted levelof exposure (<10 W/kg over any 10 g of tissue), we estimate that ˜190μW/mm² can be transferred. The low average absorption (<0.04 W/kg foradult humans) and highly localized distribution suggests that the powertransfer is unlikely to have a meaningful impact on core bodytemperatures.

The power transfer configurations evaluated here suggest thatmillimeter-scale electronics can be operated at distances sufficient forinsertion at nearly arbitrary locations in the body. Integration withexisting MEMS, logic units, sensors, light sources, and other componentswill yield many other capabilities, such as wireless communication ormechanical actuation. Applications that may emerge on this scale includedistributed sensors for the heart, deep-brain neural probes, andlocomotive transports in the bloodstream.

Materials and Methods

Numerical methods. The fields shown in FIGS. 9A-9E of the main text werecalculated from the spectral components of an in-plane source currentdensity Js (kx, ky) using the dyadic Green's function method. Thismethod reduces to a simple transfer function because the plane-wavecomponents are eigenfunctions of propagation in the multilayerstructure. At each depth z, for example, we apply a dyad G_(H) (k_(x),k_(y), z) to calculate the magnetic field H (k_(x), k_(y), =G_(H)(k_(x),k_(y), z)Jc(k_(x), k_(y)). An inverse Fourier transform yields thefields at each depth. The fields in FIGS. 10A and 10B of the main textwere calculated using a commercial electromagnetic simulator (CST StudioSuite, CST). The slot-array structure was placed above a tissue multi-layer (1 cm air gap, 4 mm skin, 4 mm fat, 4 mm muscle, 16 mm bone, 144mm heart) and the fields calculated by a time-domain solver. A similarsimulation setup—with tissue layers simplified to a slab with thedielectric properties of the liquid solution (0.5% saline)—was used toobtain the simulation curve in FIG. 2E. The port phases werereconfigured at each depth to maximize power transfer.

Theoretical power transfer calculations. The theory curve in FIG. 10Ewas calculated by considering optimal power transfer to a 2-mm coil inan air-muscle half-space at varying distances (2 to 10 cm, with a 1 cmair gap). At each depth, the optimal source Js was solved, from which|κ|/Γ_(S) was calculated from the resultant fields Es and Bs. Theremaining parameter Γ_(C) was experimentally estimated using theefficiency calibration procedure. Theory curves in FIG. 12 were obtainedin a similar manner except that an analytical solution for the fieldsdue to a single loop of wire in homogenous heart tissue (via boundaryconditions of its spherical harmonic components) was used to estimateΓ_(C). Experimental estimation was not possible across the entirefrequency range because of numerical instabilities when Γ_(C)<<1.

FIG. 12 shows efficiency of power transfer to a 2 mm diameter coil at arange of 5 cm as a function of the operating frequency. Theoreticalefficiency generated by solving for the optimal η for a multilayer modelof the chest wall (solid lines). Power transfer with the z component ofthe magnetic field (Theory z) is advantageous at low-frequencies wherethe receiver is in the near-field; in the mid-field, the transverse xcomponent of the magnetic field is dominant (Theory x). Measuredend-to-end efficiencies for six coil-based sources, designed to operateat points across the frequency range 10 MHz to 4 GHz, for the porcinebrain configuration. The error bars show fluctuation in efficiency dueto circuit nonlinearities as the receiver end power level varies from˜10-40 μW. Peak end-to-end efficiency (dot) of the slot-array structureand power transfer efficiency, calibrated for circuit inefficiencies(black square). The coil's self-resonance frequency occurs at about 3.2GHz beyond which higher order modes of the coil may contribute (shadedarea).

Electromagnetic region of operation. The frequency of operation wasselected to maximize the efficiency of power transfer to a 2-mm diametercoil at a range of 5 cm (with a 1 cm air gap). Theoretical efficiencyversus frequency curves (FIG. 12) were generated by solving for theoptimal 11 in a multilayer model of tissue (1 cm air gap, 4 mm skin, 4mm fat, 4 mm muscle, 16 mm bone, co heart) across a wide frequency range(10 MHz to 4 GHz) for coils oriented in the x and z directions, wherethe upper limit is selected to be about the self-resonance frequency ofthe coil. Coil losses were calculated using an analytical model for aloop of wire embedded in uniform tissue and impedance matching performedwith the constraint Q<10 where Q is the quality factor. Using the Dehyedispersion model for each tissue type, the peak efficiency was found tooccur at 1.6 GHz. Since each efficiency is the maximum that can beobtained by any implementation of the source, we conclude that this isthe optimal frequency of operation.

To experimentally validate the result in a complex tissue structure, wedesigned six coil-based source and receiver structures operating in thefrequency range (13, 102, 416, 950, 1560, and 2280 MHz). Eachsource-receiver pair represents a “best effort” attempt to achieveefficient power transfer in their respective electromagnetic regions.FIG. 12 shows that the measured end-to-end efficiencies replicate thegeneral shape of the curve. The peak realized efficiency is observed at1.6 GHz among the coil-based sources, with about a factor of 3enhancement provided by field focusing with the array structure.Calibrating for circuit inefficiencies, the experimental efficiency ofpower transfer is found to be within 40% of theory, although thecomposition of the intermediate tissue is highly dissimilar.

Probe trajectory visualization. A power measurement device was attachedto the end of a fiber optic cable and submerged in a liquid solution(0.5% saline). A custom-built 3D positioner (LEGO Mindstorms) moved thedevice in a “S” shaped trajectory. A photo sequence was obtained in adark room with ½ s exposure every 5 s while the device was in motion.The entire “S” shaped path was completed in 20 min. The composite imagewas created by thresholding the brightness of each image andsuperimposing the result.

Tissue imaging. Magnetic resonance imaging (MRI) of porcine tissue wasperformed at the Stanford Magnetic Resonance Systems Research Lab(MRSRL). A T2-weighted spin-echo pulse sequence was used for the heartand chest; T2-weighted fast spin-echo was used for the head.Reconstruction was performed with the OsiriX software package.

Thermal imaging. Explanted porcine loin tissue, brought approximately toroom temperature, was used for the heating experiment. The slot-arraysource was placed 1 cm above tissue and configured to couple 500 mW intotissue with uniform phase settings. Air flow between the source and thetissue was minimized by using a foam spacer. The source was brieflyremoved (<3 s) at each time point to allow infrared imaging of thetissue surface (FLIR Thermal Imaging i7). An experimental control,extracted from the same tissue piece, was placed in the field of view ofthe camera. An estimate of the overall change in temperature wasobtained by observing the hot spot and subtracting uniformly from themeasured change in background temperature. Infrared imaging of thesurface of the metal plate showed no detectable heating.

Safety considerations. We refer to IEEE guidelines for safetythresholds, although it should be noted that they are not intended formedical applications. Since our scheme relies on the coupling of bothevanescent and radiative components into tissue, thresholds based onincident power densities are not adequate. Relevant measures are insteadprovided by the specific absorption rate (SAR), defined as the powerloss integral over a reference volume. The standard includes tworelevant SAR thresholds: one averaged over the whole body (<0.4 W/kg)and another for partial body exposure (<10 W/kg, averaged over anycube-shaped 10 g of tissue). Our scheme is clearly compliant to thewhole body average—when coupling <2 W into the body, the SAR about anorder of magnitude less than the threshold (<0.04 W/kg) for a typicaladult human (60 kg).

Determining the partial body exposure requires estimation of the spatialdistribution of absorbed energy. Following recommended procedures, SARwas calculated from the initial slope of the temperature change curveSAR=C_(p)(ΔT/Δt)|_(t=0) where C_(p) is the specific heat capacity and ΔTthe change in temperature over an exposure duration of Δt, selected suchthat thermal conduction and convection effects are negligible. Forporcine tissue, the heat capacity is approximately C_(p)=3140 J/kg K,yielding an estimated SAR of 3.9 W/kg. Although the output power iscomparable to cell phones, the measured absorption is greater (>2 W/kg)because, unlike cell phones, almost all of the power is coupled intotissue. The SAR nevertheless remains well below the 10 W/kg limit forcontrolled environments. As a cross-check, power loss densities in amultilayer tissue structure were also numerically calculated using acommercial solver (CST Studio Suite, CST). When a total of 500 mW isdissipated in tissue, the maximum SAR, averaged over 10 g of tissue (orapproximately cubes of side length 2.15 cm), was found to be 3.5 W/kg.Based on these values, we estimate that we can couple up to 1.43 W intotissue before the 10 W/kg threshold is exceeded.

Source design. The slot-array source realization consists of a patternedmetal structure excited at four ports with RF signals of controlledphases (FIGS. 15A, 15B, and FIG. 10A). The patterned metal plate wasfabricated on a 1.6 mm FR4 substrate with feed and pattern copperlayers. Semi-rigid coaxial cables were used to connect each excitationport to the control board. A RF signal at 1.6 GHz was brought from thesignal generator to the control board and then separated into foursignals using a Wilkinson power divider. Following power division, thesignals are connected to parallel stages for variable attenuation, phaseshifting, and amplification. The phase shifters are voltage controlledand adjusted using a NI 9264 compact DAQ module (National Instruments)via LabVIEW (National Instruments).

FIGS. 15A and 15B show slot-array control components. FIG. 15A shows aschematic of excitation and control of the slot-array realization of thepower transfer source. When feedback from the device is available,independent control of the phase shifters enables dynamic adjustment ofthe interference pattern in tissue. FIG. 15B shows a slot-arrayimplementation consisting of two circuit boards containing the powerdivider and four parallel phase shifters and power amplifiers.

Probe design. The probe consists of a receiving coil, rectifier, chargepump, flash control integrated circuit, and an LED (FIGS. 14A and 14B).The coil was wound with an inner diameter of 2 mm using copper wire (200um diameter), with variable number of turns (between 1 to 15) dependingon the design frequency. For the rectifier circuit, two Schottky diodes(Skyworks SMS7630 series) and two 10 nF capacitors were arranged in acharge pump configuration. At low frequencies, an additional capacitorwas used in order to match the impedance of the coil and the rectifier.A charge pump and flash control integrated circuit was placed after therectifier for up-converting the rectified voltage to the 2.0 V necessaryto drive the LED. Charge was stored on a discrete 4.7 μF capacitor, anddischarged through the LED (ROHM PICOLED) when a specified thresholdvoltage was reached. The entire probe was encapsulated in epoxy,connected to a fiber optic for power information readout. The minimumpower to operate the device is about 10 μW. All components were bondedto a printed circuit board of Rogers 4350 substrate.

FIGS. 14A and 14B show circuit schematics of the power measurementprobe. FIG. 14A shows a lumped circuit model of the receiver used tomeasure power transfer. An AC voltage V_(C) is generated across the coilby the source fields. The rectifier circuitry performs AC to DC powerconversion the power flowing is encoded into the flashing frequency ofthe LED though the pulse control unit. FIG. 14B shows an equivalentcircuit at the nth reference power level. The non-linear properties ofthe rectifier and pulse control unit enables the unknown parameter R_(C)to be estimated by characterizing the circuit at two reference flashingfrequencies.

Efficiency calibration. The efficiency of power transfer wasexperimentally measured with the probe. In order to calibrate for theintermediate circuit losses, which are substantial (>50%) in the design,the non-linear behavior of the circuitry was exploited to form a set oflinearly independent equations from which unknown coil parameters can besolved. This system of equations was obtained by considering the circuitin FIG. 13A. Here, we represent the coil by an inductance L_(c), inseries with a resistance R_(c) and a voltage source V_(R) (originatingfrom the induced emf), where the latter two parameters are unknown. Fromelectromagnetic theory, the resistance can be identified asR_(C)=1/|I_(C)|²∫d³r Imε(r)|E_(C)(r)|². This quantity is identical tothe coupled-mode parameter re, aside from a normalizing constant, andcontains the integral appearing in the denominator of Eq. 1.

We characterized the non-linear circuit at two reference pulse rates, r₁and r₂, by finding the corresponding pseudo-impedancesZ_(R,n):=V_(R,n)/I_(R,n). In general, the Z_(R,n) are not equal becauseof non-linearities in particular circuit components. The correspondingamplitudes I_(R,n) and V_(R,n) were found through harmonic balanceanalysis using commercial circuit design tools (Advanced Design Systems,Agilent).

From the reduced model in FIG. 13B, we obtained the following equationsby a straightforward application of Kirchoffs voltage law:V_(C,n)=(R_(C)+Z_(R,n))I_(C,n) for n=1, 2. The two equations arelinearly independent if the reference rates correspond to operation in anon-linear region of the circuit.

For each powering configuration, an additional equation was obtained byadjusting the power level at the source until the observed pulse rateequals r₁ and r₂. Recording the two reference levels P_(s,1) andP_(s,2), an additional equation V_(C,2)/V_(C,1)=√{square root over(P_(S,2)/P_(S,1))} was obtained, permitting the unknown parameters VC,nand RC to be easily solved. The efficiency was then found by directlycalculating η=|V_(C,n)|²/8R_(C)P_(S,n).

Cardiac pacing. Adult New Zealand White Rabbits (3.0-4.0 kg) were usedin the cardiac pacing study. All animals were housed in individual cagesin the large animal facility. They were allowed to acclimatize to theholding facility for five to seven days before the procedure wasconducted. All animals were given access to food pellets, hay and waterad libitum and maintained on a 12 hour light-dark cycle (lights on at7:00 AM). All animals were fed a normal diet throughout the experimentalperiod.

The surgical site was shaved using clippers followed by a surgical scrubalternating between alcohol and Betadine. A gauze sponge was used forscrubbing. All through the procedure, draping of the surgical area wasdone appropriately. The operative procedure was performed with aseptictechnique. Surgery was performed under general anesthesia (ketamine; 35mg/kg, xylazine; 5 mg/kg, i.m. injection) using sterile technique.Antibiotics (cephazolin 20-30 mg/kg) were administered intravenously tothe animal prior to surgery. Anesthesia was maintained with endotrachealintubation and inhaled isoflurane (˜2.5-3.0%).

Once confirmed that the animal was completed sedated, a surgicalincision was made. A vertical midline incision was made and sternotomywas performed using aseptic conditions. The pericardium was excised toexpose the heart. Once the heart was in position, a 2 mm pacemakerdevice was inserted into the apex on the heart. The total procedure timewas 15 minutes.

The rabbits were sacrificed using an overdose of sodium pentobarbital(200 mg/kg, i.v. injection). Prior to that, Ketamine 20-40 mg/kg SQ andXylazine 2-5 mg/kg/SQ was administered to put the animal to sleep. Allprocedures were approved by the Animal Care and Use Committee ofStanford University.

Efficiency of Power Transfer.

The expression for efficiency in Eq. 2 above can be derived usingcoupled mode theory. In this formalism, the exchange of energy betweenthe source and receiver is described by the equations

{dot over (a)} _(S)(t)=(iω _(S)−Γ_(S))a _(C)(t)+κa _(S)(t)

{dot over (a)} _(C)(t)=(iω _(C)−Γ_(C)−Γ_(L))a _(C)(t)+κa _(S)(t)

where a_(n) are amplitudes normalized such that |a_(n)|² corresponds tothe energy in the structure, Γ_(n) the intrinsic decay rates, Γ_(L) therate of work extraction by the load on the receiver, and κ the couplingcoefficient. It may be advantageous to operate with the source andreceiver in resonance ω=ω_(S)=Ω_(C). The efficiency of power transfer isdefined as

$\eta^{\prime} = {\frac{\Gamma_{L}{a_{S}}^{2}}{{\Gamma_{S}{a_{S}}^{2}} + {\left( {\Gamma_{C} + \Gamma_{C}} \right){a_{S}}^{2}} + {{Re}\left( {\kappa \; a_{S}^{\star}a_{C}} \right)}}.}$

In the limit of weak coupling |κ|²/Γ_(S)Γ_(C)<<1, the expression reducesto

$\eta^{\prime} = {\frac{{\kappa }^{2}}{\Gamma_{S}\Gamma_{C}}\frac{\Gamma_{C}\Gamma_{L}}{\left( {1 + {\Gamma_{C}\text{/}\Gamma_{L}}} \right)^{2}}}$

which is the product of two efficiencies. The left hand factor can beunderstood as the efficiency of power transfer to the coil in absence ofthe load. The right-hand factor corresponds to the efficiency of powerextraction by the load—this factor is maximized when theimpedance-matching condition Γ_(C)=Γ_(L) is satisfied and, as aconsequence of the maximum power transfer theorem, is at most 25%. Fromstandard power arguments, it can be shown that the left-hand efficiencyis given by

$\begin{matrix}{\frac{{\kappa }^{2}}{\Gamma_{S}\Gamma_{C}} = \frac{{{\int{d^{3}{{rB}_{S}^{*} \cdot M_{C}}}}}^{2}}{\left\lbrack {{\int{d^{3}r\mspace{11mu} {Im}}} \in {(\omega){E_{S}}^{2}}} \right\rbrack \left\lbrack {{\int{d^{3}r\mspace{11mu} {Im}}} \in {(\omega){E_{C}}^{2}}} \right\rbrack}} & ({S1})\end{matrix}$

which is the efficiency in Eq. 2. Equivalent expressions can be obtainedusing other models for coupled electrical systems, such as a two-portlumped element network.

Penetration of time-varying fields in tissue. Although electromagneticwaves varying in time at high frequencies (>100 MHz) are associated withhigh absorption in tissue, optimal transfer in FIG. 9A is found to occurin the low-gigahertz range. To understand this result, we consider thepenetration of a plane wave H_(x)(z, t)=H₀exp(i(kz−ωt)) into a z<0tissue half-space as the frequency ω varies. The depth at which thefield extends into tissue is described by the skin depth δ:=1/Im(k)where k=ω√{square root over (με)} is the wavenumber. Dispersion intissue can be described by Debye relaxationε/ε₀=ε_(∞)+Δε/(1−iωτ_(D))+iσ/(ωϵε₀) where σ is the conductivity, τ_(D)the characteristic relaxation time of the medium, ε_(∞) the permittivityin the high frequency limit, and Δε:=ε_(S)−ε_(∞) (ε_(S) the staticpermittivity). The model is valid in the regime ωτ_(D)<<1, in which casethe permittivity can also be approximated as

ε/ε₀≈ε_(S)[1+i(ωτ_(D)Δε/ε_(S)+1/ωτ_(E))]  (S2)

where τ_(E):=ε₀ε_(S)/σ is the electric time constant. The relativesignificance of ωτ_(E) and 1/ωτ_(D) determines the imaginary term'scharacteristic dependence on ω:1. At low frequencies when ωτ_(E)<<1, the permittivity reduces toε/ε₀≈iε_(S)/ωτ_(E). The wavenumber is given by k≈ω√{square root over(μ₀/2(1+i))}, from which one obtains the usual skin depth for conductorsδ≈√{square root over (2/ωμ₀σ)}.2. When ωτ_(E)>>1 but ωτ_(E)<<1/ωτ_(D), the τ_(D) term in Eq. S2 can beneglected such that δ=2τ_(E)/√{square root over (μ₀ε₀ε_(S))}.3. At high frequencies when ωτ_(E)>>1/ωτ_(D), we approximate thepermittivity as ε/ε₀≈ε_(S)(1+iωτ_(D) Δε/ε_(S)). The skin depth is thengiven by δ≈2ε_(S)/(ω²τ_(D)Δε√{square root over (μ₀ε₀ε_(S))}. Theparameter 1/τ_(E) and the geometric mean of 1/τ_(E) and 1/τ_(D)demarcate three frequency regimes

$\delta \propto {\begin{Bmatrix}{{1\text{/}\sqrt{\omega}},} & {\omega {\operatorname{<<}1}\text{/}\tau_{E}} \\{{Ο(1)},} & {1\text{/}\tau_{E}{\operatorname{<<}\omega}{\operatorname{<<}1}\text{/}\sqrt{\tau_{D}\tau_{E}}} \\{{1\text{/}\omega^{2}},} & {\omega\operatorname{>>}{1\text{/}\sqrt{\tau_{D}\tau_{E}}}}\end{Bmatrix}.}$

Contrary to the notion that losses consistently increase with frequency,we find that there exists an intermediate range of frequencies acrosswhich the penetration is approximately constant. This behavior occurswhen the typical time scale of an amplitude variation is much shorterthan τ_(E) but substantially longer than τ_(D). If one operates in thisregime, the penetration is expected to be much greater than that naivelyextrapolated from a low-frequency conductor model of tissue. For themuscle tissue parameters, we calculate the range to be approximatelybetween 690 MHz and 2.2 GHz, which is consistent with our implementationof midfield power transfer.

State-of-the art integrated electronics. To illustrate the range ofapplications available with performance characteristics reported in themain paper, FIG. 16 describes the power requirements of selectedstate-of-the-art integrated circuits (ICs). The table is not exhaustive,but is representative of existing solid-state circuit capabilities inthe microwatt power regime. With the exception of (31), all devices arecurrently powered with either wire tethers or large (>2 cm) near-fieldcoils. For stimulation, a local field sensing IC was developed byMedtronic to enable closed-loop neurostimulation (32). As an alternativeto electrical stimulation, an optogenetic stimulator consuming 400 μWwas designed, with possibility of lower requirements by using moreefficient LEDs or opsins (33). Using a more advanced process node andsub-threshold design techniques enabled a reduced power consumption of0.73 μW, as demonstrated for neural recording (34). Less than 100 μW wasrequired for 100-channel neural recording. A pacemaker IC, developed bySt. Jude Medical, contains amplifiers, filters, ADCs, battery managementsystem, voltage multipliers, high voltage pulse generators, programmablelogic, and timing control; and consumes only 8 μW (24). A wide dynamicrange bio-impedance sensor was also shown to extract QRS features forthe detection of ventricular fibrillation (35).

ICs developed for monitoring physiological processes include afluorimeter for continuous glucose monitoring (36), a cubic-millimeterintraocular pressure sensor (37), and a temperature sensor with anaccuracy of ±0.15° C. (38). These sensors require energy in the range ofnJ to μJ per measurement. For imaging, a sensor consuming only 3.4 μJper frame of 256×256 pixels with 8 bits per pixel has been demonstrated(39). An implantable device capable of locomotion in fluid has also beendeveloped (31). Wireless communication allow remote control andnon-invasive readout of these devices. In (31), the data receiverconsumes 0.5 pJ per bit. For the reverse link (from the device to theexternal source), the power consumption depends on the range, varyingbetween pJ to nJ per bit.

While the present disclosure (which includes the attachments) isamenable to various modifications and alternative forms, specificsthereof have been shown by way of example in the drawings and will bedescribed in further detail. It should be understood that the intentionis not to limit the disclosure to the particular embodiments and/orapplications described. Various embodiments described above and shown inthe figures and attachments may be implemented together and/or in othermanners. One or more of the items depicted in the drawings/figures canalso be implemented in a more separated or integrated manner, as isuseful in accordance with particular applications.

As for additional details pertinent to the present invention, materialsand manufacturing techniques may be employed as within the level ofthose with skill in the relevant art. The same may hold true withrespect to method-based aspects of the invention in terms of additionalacts commonly or logically employed. Also, it is contemplated that anyoptional feature of the inventive variations described may be set forthand claimed independently, or in combination with any one or more of thefeatures described herein. Likewise, reference to a singular item,includes the possibility that there are plural of the same itemspresent. More specifically, as used herein and in the appended claims,the singular forms “a,” “and,” “said,” and “the” include pluralreferents unless the context clearly dictates otherwise. It is furthernoted that the claims may be drafted to exclude any optional element. Assuch, this statement is intended to serve as antecedent basis for use ofsuch exclusive terminology as “solely,” “only” and the like inconnection with the recitation of claim elements, or use of a “negative”limitation. Unless defined otherwise herein, all technical andscientific terms used herein have the same meaning as commonlyunderstood by one of ordinary skill in the art to which this inventionbelongs. The breadth of the present invention is not to be limited bythe subject specification, but rather only by the plain meaning of theclaim terms employed.

1. (canceled)
 2. An apparatus configured for use in transferring powerwirelessly through tissue, the apparatus comprising: multiple excitableelectromagnetic structures provided externally to a patient body andconfigured to generate, in response to respective first excitationsignals, a non-stationary field inside the patient body; and acontroller configured to change a power transfer characteristic of thenon-stationary field by changing one or more of the first excitationsignals provided to the electromagnetic structures.
 3. The apparatus ofclaim 2, wherein the controller is configured to change the powertransfer characteristic of the non-stationary field inside the patientbody by changing a frequency, amplitude, or phase characteristic of atleast one of the first excitation signals.
 4. The apparatus of claim 2,wherein the controller is configured to change a travel direction of thenon-stationary field inside the patient body by changing the one or moreof the first excitation signals.
 5. The apparatus of claim 2, whereinthe controller is configured to change a magnitude of power transmittedto a receiver device inside the patient body by changing the one or moreof the first excitation signals.
 6. The apparatus of claim 2, whereinthe multiple excitable electromagnetic structures are configured togenerate an evanescent field outside of the patient body that influencesa propagation characteristic of the non-stationary field inside thepatient body.
 7. The apparatus of claim 2, further comprising multipleindependent feed ports configured to excited respective ones of themultiple excitable electromagnetic structures to generate thenon-stationary field inside the patient body.
 8. The apparatus of claim2, wherein the multiple excitable electromagnetic structures are furtherconfigured to generate a steerable field inside the patient body basedon adjustable characteristics of the first excitation signals.
 9. Theapparatus of claim 2, wherein the controller is configured to control asignal generator to provide the first excitation signals to theelectromagnetic structures, wherein the first excitation signalscomprise signals between about 300 MHz and 3000 MHz.
 10. The apparatusof claim 2, wherein the controller is configured to update an outputpower of the apparatus in response to information received from animplanted electrostimulation device about a quantity or quality of powerreceived from the apparatus.
 11. The apparatus of claim 2, wherein thecontroller is configured to update a focus of the non-stationary fieldin the patient body in response to information received from animplanted electrostimulation device about a quantity or quality of powerreceived from the apparatus.
 12. The apparatus of claim 2, wherein thecontroller is configured to update a focus of the non-stationary fieldin the patient body to concentrate the field in a direction of aparticular one of multiple implanted receiver devices.
 13. A method forwirelessly providing power to an implanted device from an externaltransmitter apparatus, the method comprising: generating firstexcitation signals; exciting multiple electromagnetic structuresconcurrently using respective ones of the first excitation signals, theelectromagnetic structures provided externally to a patient body andgenerating, in response to the first excitation signals, anon-stationary field inside the patient body; and updating a powertransfer characteristic of the non-stationary field by changing a signalcharacteristic of one or more of the first excitation signals.
 14. Themethod of claim 13, wherein the updating the power transfercharacteristic includes changing a focus of the non-stationary field inthe patient body.
 15. The method of claim 13, wherein the changing asignal characteristic of one or more of the first excitation signalsincludes changing one or more of an excitation signal frequency,amplitude, or phase.
 16. The method of claim 13, wherein the updatingthe power transfer characteristic includes changing a magnitude of powertransmitted to the implanted device.
 17. The method of claim 13, furthercomprising, in response to the exciting the multiple electromagneticstructures concurrently using respective ones of the first excitationsignals, generating an evanescent field outside of the patient body thatinfluences a propagation characteristic of the non-stationary fieldinside the patient body.
 18. A transmitter configured for use intransferring power wirelessly through tissue, the transmittercomprising: means for receiving first excitation signals and in responseproviding a non-stationary field inside the patient body; and means forupdating a power transfer characteristic of the non-stationary field bychanging a signal characteristic of one or more of the first excitationsignals.
 19. The transmitter of claim 18, wherein the means for updatingthe power transfer characteristic of the non-stationary field includesmeans for changing a focus of the non-stationary field in the patientbody.
 20. The transmitter of claim 18, wherein the means for updatingthe power transfer characteristic of the non-stationary field includesmeans for changing a frequency, amplitude, or phase of one or more ofthe first excitation signals.
 21. The transmitter of claim 18, whereinthe means for receiving the first excitation signals includes means forchanging, in response to the first excitation signals, an evanescentfield outside of the patient body to thereby influence a propagationcharacteristic of the non-stationary field inside the patient body.